Tissue engineering is an attempt to regenerate a defect in tissue that is larger than the unaided body can regenerate on its own. In most cases tissue engineering requires technology for the creation of three components: implants (often referred to as scaffolds), cells and growth factors. Tissue regeneration not only requires the infusion of cells specific to the function of the organ but also vasculature and often connective tissue. Growth factors can aid in the performance of concentrated tissue precursor cells or the recruitment of reparative host tissue. Implants or scaffolds are often required to provide guidance to stem cells and/or invading host tissue, vasculature and connective tissue. Implants may be designed to match a defect in a patient's tissue. The shape of the implant may be determined by first measuring the defective area or volume within the patient. The implant may then be designed by, for example, computer aided design (CAD) in light of the measured defective area or volume. The implant may then be manufactured.
Factors to take into account when designing and manufacturing implants include adequate geometry to provide a proper fit: (a) the external surface fit of the implant into the defect site, and (b) the porous space within an implant to guide the initial infusion of tissue, vasculature, and connective tissue. If the walls between porous spaces of the implant or scaffold are too thick they may not resorb, and thereby become a barrier to remodeling. If the materials degrade, their byproducts need to be non-toxic and easily metabolized so that they do not prevent tissue regeneration or remodeling.
Functional geometrical features of a scaffold may be designed to affect cell attachment, proliferation, or maturation. Surface features that interact directly with cells include scaffold roughness and porosity. Rough, porous structures may facilitate cell loading, neotissue growth, and host tissue ingrowth. The designer may manipulate porous geometry to control both the mechanical properties of the whole implant as well as the pore space's porosity, tortuosity, permeability, and total pore volume. Many tissue engineering scaffolds may require pores that range between 200 and 1600 micrometers with varying surface features, such as the shape of the pore opening, in the order of 50-500 micrometers. Conventionally, these features may have been obtained by the inclusion of particles such as tricalcium phosphate crystals into the resin from which the scaffold would be manufactured. However, concerns may arise as to the resorbability of the crystals in the host's body.
Another important geometrical feature may be oblique orientation of pore structures in order for the host tissue to not encounter a wall or barrier in the scaffold, which is more likely when pore structures are built orthogonally than when pores or channels are oriented towards host tissue. The implant designer may want to orient pores channels within a scaffold so that they open toward the host tissue thereby facilitating growth of new tissue into the implant and active incorporation of the implant into the host tissue.
Additive manufacturing of implants or scaffolds with these mechanical and geometrical features requires relatively high accuracy levels. For example, accurate rendering makes it more likely that complex internal pore structures such as those described above and other can be created. Stereolithography is described by Paul Jacobs in: Rapid Prototyping & Manufacturing: Fundamentals of StereoLithography by Paul F. Jacobs (Jan. 15, 1992), and Stereolithography & Other RP&M Technologies: From Rapid Prototyping to Rapid Tooling by Paul F. Jacobs (Jan. 1, 1996).
Additional factors to take into account when designing and manufacturing implants or scaffolds are adequate strength and stiffness for the part to handle and transmit mechanical stress. In some cases, strength and stiffness must be weighed against the need for the implant or scaffold to be resorbable or capable of breaking down in the host's body. Manipulation of the polymer's molecular weight often adjusts both the rate and extent of resorption levels in vitro as well as in vivo versus strength of the implant, with higher molecular weights often being stronger and lower molecular weights often being more resorbable. However, post-curing handling of low molecular weight scaffolds or implants could be problematic and thus the ideal rendering method would minimize any post-curing handling necessary.
While stereolithographic rendering of implants and scaffolds has been demonstrated, limitations in the commercially available devices has thus far resulted in relatively low accuracy levels.
For example, accuracy and resolution of conventional stereolithographic rendering devices may not allow the devices to produce scaffold or implant surface features such as pores and pore openings at the low end of the optimum geometry scale. While conventional stereolithographic rendering devices may be able to produce orthogonally oriented pore structures in implants and scaffolds, they may not be able to provide sufficient resolution to produce obliquely oriented pores. Moreover, stereolithographic rendering may also have various other limitations in the context of manufacturing of implants or scaffolds. For example, conventional stereolithography devices use a laser to polymerize layers. The laser points downward at the top of a vat of liquid polymer. An elevator sits inside the vat and pulls the part downward as it is rendered, layer by layer. The drawing speed is typically not fast enough to simultaneously draw all pixels in the layer, which may make it difficult to control pixel to pixel crosslinking within the layer and/or over-curing or stitching between layers as the implant or scaffold is rendered.
Also, conventional stereolithography devices may not provide a way to modulate the amount of energy at one spot versus another within a layer to, for example, control the depth of polymerization and level or strength of over-curing. Controlling the depth of polymerization as well as the level or strength of curing is critical.
Control of resolution in chain length dependent propagation with continuous digital light processing (“cDLP”) as in many other forms of photo-initiated additive manufacturing is essential to render useful and accurate parts. Several important technological aspects that allow for highly accurate additive manufacturing are (i) accurate delivery of light, (ii) good control of the wavelength and amount of energy in that light, and (iii) a build surface that can be moved into an appropriate position to form each layer and have it bind (i.e. laminate) with the previously built layer.
Normally, in cDLP manufacturing, light inhibiting agents, known as dyes (also referred to as light attenuators), are introduced to a polymer mixture in order to limit the wavelengths of light that activate a photo-initiator as a means to control the depth of curing, or the z-axis resolution. These dyes are numerous, however, the selection of biocompatible ones are much less. Furthermore, dyes that are effective against ultraviolet transmission and USP grade are even harder to come by. One dye that fits these requirements is titanium dioxide. Like many ceramics, it is biocompatible, stable, and small in particle size, making it ideal for use in photo-initiated polymer mixtures. However, in its effectiveness, there is an inevitable downside, while titanium dioxide is a strong ultraviolet absorber, it also has strong scattering properties, which leads to inadvertent curing, thereby decreasing resolution in the xy-plane and potentially z-plane as well. This phenomenon is called “dark-cure.” Identification of dyes or other biocompatible agents that function together with titanium dioxide, as well as on their own to produce resorbable, biocompatible tissue engineering scaffolds with desired physical, biological and chemical properties are needed.